In 1977, Gruentzig underwent the first case of percutaneous transluminal coronary angioplasty (PTCA), breaking through the therapy situation of drugs and surgery and creating a new era of interventional cardiology. Since the development of interventional therapy of coronary heart diseases, a percutaneous transluminal coronary angioplasty (PTCA) era, a bare metal stent (BMS) era and a drug-eluting stent (DES) era have been experienced. The vascular restenosis rate is reduced to below 10% by a drug stent from 50% of balloon dilatation alone and 20 to 30% of a bare metal stent, especially when small vascular diseases are suffered or a lesion period is longer. The advantages of the drug stent are very obvious.
The coronary stents commonly used in clinic falls into two categories: bare metal stents and drug-eluting stents. Currently, the market share of the drug-eluting stents reaches 95% in China, but in foreign countries, the market share of the bare metal stents is still 30% to 50%. This is because that although the restenosis and revascularization rates can be reduced by the DES, the existing polymer carrier drug stents still have some limitations, mainly showing late and very late stent thrombosis problems, delayed endothelial healing and late catch-up of lumen loss, and the main reason is polymer carrier-induced inflammation. The problems and the effective means for solving the problems have been widely debated in the field of international research. One research field is to develop a fully biodegradable polymer coating drug-loaded stent, and the other development field is to avoid the use of a polymer coating. i.e., a carrier free drug stent. However, since a substrate material belongs to a permanent implant, its long-term potential risks still exist.
The therapy method of infant congenital vascular stenosis (coarctation of the aorta and pulmonary stenosis) includes surgery, balloon angioplasty and stent implantation. Although surgery is a good method, it is a thoracotomy and has a big trauma; meanwhile it is difficult to solve the pulmonary branch stenosis and postoperative restenosis. The balloon angioplasty and the stent implantation, which are safe transcatheter interventional therapy methods, have advantages of less trauma and shorter hospital stay, etc.; but the balloon angioplasty leads to a higher incidence of complications, especially for babies; therefore, the stent implantation is proved to be a better choice. However, infants have further growth and development characteristics, and the non-absorbable stents implanted can cause restenosis in the late period of vascular growth; although the stent diameter can be consistent with the vascular growth by a re-expansion method, adult stents cannot be implanted into blood vessels of infants because blood vessels of infants are thinner.
Currently, the bioabsorbable vascular stent has become a research focus, and has the advantages that the ordinary stents do not have: further growth of blood vessels and the follow-up vascular surgical therapy cannot be hindered; after the bioabsorbable stent is completely absorbed by the human body, narrow blood vessels will be restored to healthy and natural normal blood vessels with a physiological vasomotor capacity; the stent can be fully absorbed until the stent disappears completely, so that the chronic injury and inflammatory reaction caused by the stent for a long term can be avoided and the late stent thrombosis is reduced and so anti-platelet drugs do not need to be taken for a long term; once the stent is completely absorbed, the stent does not have the long-term potential adverse effects on the blood vessels without increasing the surgery difficulty of re-PCI or surgical revascularization, which has a great significance, especially for the blood vessels of children in the period of growth and development.
The bioabsorbable stents mainly comprise a polymer-based bioabsorbable stent and a metal-based bioabsorbable stent. But the former has unsatisfactory biomechanical properties, and simultaneously the complexity of such stent release process is much higher than that of the conventional balloon dilatation metal stent. The latter mainly includes a magnesium alloy stent and an iron stent at present. The magnesium alloy stent cannot play an effective supporting role before revascularization due to its too fast corrosion rate; therefore, the development focus of the magnesium alloy stent lies in how to reduce the corrosion rate thereof. Pure iron applied to the bioabsorbable stents has the main disadvantages of low mechanical properties and too slow corrosion rate. In the prior art, a composite coating containing strontium or calcium, or both is prepared on the surface of bioabsorbable metal materials such as pure iron by physical vapor deposition to accelerate and control the corrosion rate of the materials. In addition, a polymer coating which can be degraded in an acidic environment is sputtered on the composite coating to further accelerate the corrosion rate of the materials. However, such a method fails to solve the problems of low mechanical properties of pure iron materials. The coating and substrate pure iron have a problem on whether both are firmly bound or not due to their non-integrated structure.
The current research mainly focuses on developing novel iron-based alloys and finding novel iron material preparation methods, or preparing an iron alloy layer on the surface of the pure iron material and modifying the pure iron material in order to solve the problems of the pure iron stents; wherein the pure iron stent is subjected to surface alloying treatment (carburizing, nitriding, and carbonitriding) to obtain a composite diffusion layer with an adjustable permeation depth, thereby improving the strength of the stent, and simultaneously accelerating the corrosion rate of the stent and shortening the absorption cycle of the stent. After the stent is subjected to surface alloying, the stent has a non-continuous diffused composite diffusion layer. By controlling the distribution, shape and depth of the diffusion layer, the yield strength and elongation can be adjusted in a wide range to achieve the strength and absorption cycle required by the stent. The composite diffusion layer comprises a solid solution that nitrogen exists in iron, and Fe4N.
The bioabsorbable metal represented by pure iron and magnesium alloy can be used for manufacturing other implantable medical devices other than for manufacturing bioabsorbable vascular stents.
The vascular stent subjecting to a surface treatment method such as nitriding, carburizing or carbonitriding and then being polished in the prior art has a composite diffusion layer with the adjustable permeation depth. An unsolved technical problem is how to optimize the structure on this basis to improve the comprehensive properties (radial strength, flexibility, fatigue resistance and corrosion rate) of the absorbable stent.
The radial strength of the stent is defined herein as a pressure required when the stent radially deforms by 10%. For a coronary stent with an outer diameter of 1.6 mm, the outer diameter is dilated to 3.0 mm by a balloon generally and then the radial strength is measured. A 316L stainless steel coronary stent has a wall thickness of about 100 μm usually, and a radial strength ranging from 110 KPa to 150 KPa; the mechanical properties of cobalt-chromium alloy is slightly better than those of 316L stainless steel, and a cobalt-chromium alloy coronary stent has a wall thickness of about 80 μm usually with a radial strength ranging from 140 KPa to 185 KPa.
The blood vessels of human body usually tend to be bent or twisted, especially a vascular lesion segment. Flexibility of the stent refers to a capacity of the stent to adapt to bent blood vessels. The better the stent flexibility is, the stronger the penetration capacity of the stent through the blood vessels is. According to the finite element analysis of the stent, a wall thickness of the stent not only is a main factor affecting the stent flexibility, but also one of key parameters reflecting the comprehensive properties of the stent. Meanwhile, the wall thickness of the stent is regarded as an independent predictive factor of late lumen loss (vascular restenosis) after vascular lesions are treated by an interventional therapy, and the evidence-based medicine agrees that the restenosis rate of a thin-walled stent is lower than that of a thick-walled stent. However, the wall thickness of the stent is reduced at the expense of the loss of radial strength of the stent; therefore the wall thickness of the stent is strictly limited by clinical requirements. The commonly used coronary stents are all permanently implantable, including bare metal stents and drug-eluting stents, in which 316L stainless steel or cobalt-based alloy are usually taken as a substrate material. However, the mechanical properties of the stent depend on the substrate material and the stent structure design, i.e., the mechanical properties of the coronary stent can not be affected by drugs basically. Under the premise of ensuring the clinical therapeutic effects, the wall thickness of the current permanently implantable coronary stent can only be reduced to 65 μm, and the stent uses cobalt-chromium alloy as the substrate material. Under the same mechanical property requirements, although the wall thickness of the iron-based coronary stent is significantly less than that of the coronary stents made from other bioabsorbable materials, the wall thickness of the bioabsorbable iron-based coronary stent obtained by the prior art can only be reduced to 90 μm or so, and does not reach the minimum wall thickness of 65 μm of the permanently implantable coronary stent. The technical problem on how to reduce the wall thickness of the bioabsorbable iron-based stent as much as possible under the premise of satisfying the mechanical properties such as elongation and radial strength and improving the corrosion rate of the stent has not been solved yet by the prior art.
The iron-based material (including pure iron, steel or other iron alloys) is subjected to a surface nitriding process such as ion nitriding, so that a denser compound layer is formed on the surface of the iron-based material generally. According to the known research results, the compound layer is formed by γ′ phase (mainly Fe4N), or by mixing γ′ phase and ε phase (composition change range Fe2-3N); wherein the γ′ phase accounts for 50 to 100% by weight. The nitrogen content of the γ′ phase is 6 wt-% or so, and the nitrogen content of the ε phase is 8 to II wt-% or so; therefore, nitrogen atoms in the compound layer have a very high concentration, and are diffused to inside of the material through high temperatures. The compound layer is easily formed on the surface of the iron-based stent after the iron-based stent is subjected to surface nitriding treatment, and can greatly increase brittleness of the material; the corrosion resistance of the compound layer is much higher than that of a pure iron substrate. Therefore it is necessary to fully remove the compound layer under the premise of ensuring the properties of the stent. When the iron-based stent is prepared by a prior method, if the plasma discharge bias is too low (below 600V), the average thickness of the compound layer is generally more than 10 μm. If the temperature of the iron-based stent is too high (above 550° C.), the local compound layer on the surface of the iron-based stent extends to inside of the material in a dendritic or flaky shape, resulting in very uneven thickness of the compound layer. In the prior art, a pure iron pipe subjected to drawing and mechanical polishing is used, and the grain boundary is in a disordered high-energy state due to work hardening; meanwhile, there are higher internal residual stress inside the pipe and more defects such as dislocations, thereby providing more express channels for diffusion of nitrogen atoms; the compound layer will extend to a deep part inside the iron pipe material along the grain boundary or a dense dislocation area to show an inward dendritic morphology. The crystal defects of the surface can be significantly reduced by fully annealing the pure iron pipe, but this is not conducive to the permeation of nitrogen ions; therefore it is difficult to solve the problem. In view of the subsequent imprecise and non-uniform surface polishing treatment (the larger the thickness is removed by polishing, the more unfavorable), a compound layer with a certain thickness or higher coverage rate is possibly remained after the surface of such stent is polished; therefore, the purpose of improving the corrosion rate of the iron-based stent may not be achieved.
After being implanted into the blood vessel, the absorbable stent must maintain sufficient mechanical properties within an initial period of time (several months or longer) to adapt to the bending shape of the blood vessel and block the collapse of lesion blood vessel, and can be gradually absorbed after the vascular remodeling is stable. If the local strain of the stent exceeds a certain limit, micro-cracks will be firstly generated on the surface of a part of supporting strut or connecting part. Due to vascular pulsating and blood flow, the metal fatigue of the stent will be gradually accumulated; at this time, the micro-cracks will gradually propagate to inside from the surface of the stent to become larger cracks damaging the stent structure until the stent is partially fractured, and it is even more necessary to prevent the propagation of the micro-cracks for the thin-walled stent. Therefore, the prior art needs to be optimized to ensure that sufficient pure iron or a low-nitrogen-content area is still reserved inside the nitrided stent substrate to reduce the risks of surface crack propagation and premature brittle failure of the stent (before the vascular remodeling is stable). Hence, the ratio of the depth of a nitrided layer to the wall thickness of the stent cannot be too large.
In order to obtain a better biological tissue compatibility, the roughness of inner and outer surfaces of the coronary stent should be reduced as much as possible. The ordinary electrochemical polishing used in the prior art does not have a good polishing effect on the inner wall of the stent, and the surface roughness can only be controlled below 0.1 μm, which cannot reach a mirror bright effect (surface roughness≦0.01 μm); the stent surface can be bright and smooth when the polishing removal amount (the difference value between wall thicknesses of a polished stent and an unpolished stent) reaches 40 μm above in the prior art; thus the inherent shortcomings of imprecision and unevenness of polishing treatment are more obvious, and especially not conducive to the quality control of thin-walled coronary stents.
An important technical problem related to this is that the prior art is difficult to be used for a thin-walled pipe (wall thickness is less than 100 μm). If the original pipe wall of the stent pipe is very thin, the nitrided layer is bound to be thinner. However, in order to achieve the desired polishing effect, the prior polishing method requires a higher polishing removal amount, the compound layer closest to the surface can not only be removed, but also a part of nitrided layer can be removed, so that the remaining nitrided layer will become very thin. The thickness uniformity of the nitrided layer is limited by the prior art, the polishing removal amount of different areas of the supporting strut are not uniform enough, and the two uniformities are superimposed together, resulting in the more non-uniform thickness of the remaining nitrided layer. If the remaining nitrided layer is too thin, the nitrided layers at some parts of the supporting strut will almost completely disappear, thereby bringing serious adverse effects. The nitrided layer can effectively improve the radial strength of the stent, which is especially critical to the thin-walled stent; if the thickness of the nitrided layer is not very uniform, the mechanical properties of the various parts of the stent are very inconsistent; thus, the radial strength of the stent will not meet the design requirements. In addition, if the nitrided layer is very thin and has a non-uniform thickness, some parts of the supporting strut are slowly corroded due to less nitrogen content; thus the stent can also not meet the design requirements. By further considering the process error in actual production, the difference between distant supporting struts or connecting parts will be more obvious, and the problems of inconsistency of the above-mentioned mechanical properties or less local nitrogen content will be more serious. Therefore, the polishing process in the prior art needs to be improved so as to adapt to the thin-walled pipe with a wall thickness of less than 100 μm.